1. Field of the Invention
This invention generally relates to the determination of the three-dimensional distribution and intensity of multiple radioisotopic sources of essentially unknown spatial and/or temporal characteristics in the presence of noise by means of appropriate measurement and analysis and, more particularly, to a radioisotopic imaging system for accurately reconstructing with a high spatial resolution three-dimensional distributions of radioactivity of the kind encountered in lesion detection in nuclear medicine.
2. Description of Related Art
Radioisotopic organ imaging is customarily carried out by (a) administering a gamma ray emitting isotope to a patient, (b) determining the radioisotopic distribution and intensity in the organ by detecting gamma rays exiting the organ, and (c) analyzing the detected gamma ray data. The detection is performed by collimator detector assemblies which have fields of view which must be taken into account.
Focussing collimators were employed initially for selectively detecting radiation emitted within an approximately cylindrical field of view extending through the focal region as specified by their point source response function (PSRF). The clinical imaging procedure consisted of moving a collimator detector assembly along a series of planar parallel lines relative to the patient and generating a raster output.
Focussing collimator systems were not well adapted to study dynamic phenomena such as cardiac function and have generally been replaced by gamma cameras which continuously detect gamma rays emitted from the whole region of interest. This was done with parallel hole or pin-hole collimators, each of which served to define a restricted field of view necessary for effective determination of the radioisotopic tissue distribution.
For three-dimensional imaging, it was necessary to view the tissue of interest at many different angles and to analyze the data with the aid of a computer using one or more standard image processing algorithms. Collimation limited the gamma rays arriving at the detectors and less than about one-thousandth of all gamma rays exiting the patient were detected. The number of detections or counts recorded in individual pixels (data elements) was variously limited by this inefficiency, by the energy resolve time of the sodium iodide (NaI) crystal detectors generally used (about 250 ns), by the practical time limit for patient immobilization (about 10-15 minutes), and by patient and personnel radiation exposure safety considerations. Statistical fluctuations were therefore quite significant.
The field of view defined by each channel in the parallel-hole collimator was a solid diverging cone specified by the channel geometry. A true pin-hole collimator had zero detection efficiency. However, in practice, the pin holes had significantly non-zero dimensions. Together with the uncertainty in specifying the exact site of the gamma ray-crystal detector interaction, the emission probability field (EPF) associated with a single gamma ray-induced crystal flash using pin-hole collimation was again a solid diverging cone.
Because of the relatively large proportion of statistical noise and the imprecision associated with the solid diverging conical emission probability fields (EPFs) defined by individual measured events, the tomographic cold lesion detection limit attainable in standard clinical nuclear imaging was about one to two cubic centimeters using a single photon emission computed tomography (SPECT) system. This limit was well above that attainable with other higher resolution imaging modalities.
Nonetheless, radioisotopic organ imaging continued to be very important as an index of physiological function as distinguished from other types of anatomical properties. In an effort to improve the resolution possible in radioisotopic imaging, positron-emission tomography (PET) was developed. A variety of different systems existed, but they were all based upon coincidence-counting the 0.511 Mev gamma ray pairs arising from positron-electron annihilation and specifying the almost straight angle (about 180.degree.) between the gamma rays of each pair. Since the generally cylindrical EPF extending about and along the path of propagation in a PET system was generally smaller than the solid diverging conical zone of earlier methods, somewhat higher resolutions were achieved. The limit on resolution in a PET system arose, among other factors, from the mean free path of the positron prior to capture in a detector, the deviation from a straight angle path of each pair of gamma rays due to the residual momentum of the positron prior to annihilation, the longitudinal uncertainty in the annihilation site, and the interaction site uncertainty in the about three cubic centimeter thick unshielded NaI detectors usually used. These relatively large, thick detectors were required to absorb the high energy gamma rays and to exclude secondary gamma rays arising from Compton scattering within the tissue of interest. Denser, more compact, solid-state, thinner detectors were also used to reduce this latter limitation. Time of flight information was used to try to reduce the emission probability field along the longitudinal direction, but even a 0.2 nanosecond resolve time involves an annihilation site longitudinal ambiguity of about six centimeters.
Mechanical detector wobbling has also been proposed as a means of improving the resolution. In general, however, cold lesions less than about one to two cubic centimeters remained undetected. In view of the specialized facilities required for the short-lived isotopes commonly used (medical cyclotrons) and the limited improvement in resolution, PET systems have not found general application.
A focussing collimator coincidence scanning (FCC) system with very small lesion detection capability (resolutions less than about 0.01 cm.sup.3) was proposed by H. Hart and S. Rudin, IEEE Trans. Biomed. Eng., BME-24,169,1977, and was based upon coincidence detection of gamma rays from isotopes which emitted more than one gamma ray in cascade. By using multiple focussing collimators with intersecting fields of view, coincidence events served to define a very small focal region. Laboratory detection of structures having a volume of less than 0.01 cm.sup.3 was reported.
The sensitivity of the FCC system was, however, very low. Clinical scan times would have been generally unacceptable on a routine basis. Dynamic imaging was not possible. The FCC approach has not been clinically applicable.
A system including a single pair of uncollimated planar detectors was proposed by M. Singh and D. Dario in IEEE Trans., Nuc. Sci., NS-31,594; 1984. Electronic collimation was substituted for material absorption collimation. This substitution greatly improved the sensitivity possible for gamma cameras. The resolution reported with a prototype system seemed to be considerably greater than one centimeter. The ultimate resolution possible appeared to be limited since the emission field defined by a two-fold (thin-thick detector) coincidence event is a relatively large hollow conical shell whose wall thickness is a function of the uncertainties in the detector interaction sites and the Compton scattering deflection angle arising from the finite energy resolution of the thin detector component of the system. See, also, L. Kaufman et al., IEEE Trans., Nuc. Sci., NS-27, 1073, 1980.
In all previously known radioisotopic systems, there has been no radioisotopic imaging system demonstrated or proposed capable of practical, clinical, dynamic, tomographic imaging on a clearly sub-cubic centimeter scale. Such a capability would be extremely important, both clinically and in research.